The present invention relates to tomography, and more particularly to correcting errors caused by extraneous radiation in tomographic systems, such as cone-beam computerized tomography (CBCT) systems, fluoroscopic tomography systems, radiographic tomography systems, laminar tomography imaging systems, and the like.
Computerized tomography (CT) involves the imaging of the internal structure of an object by collecting several projection images (“radiographic projections”) in a single scan operation (“scan”), and is widely used in the medical field to view the internal structure of selected portions of the human body. Typically, several two-dimensional projections are made of the object, and a three-dimensional representation of the object is constructed from the projections using various tomographic reconstruction methods. From the three-dimensional image, conventional CT slices through the object can be generated. The two-dimensional projections are typically created by transmitting radiation from a “point source” through the object, which will absorb some of the radiation based on its size, density, and atomic composition, and collecting the non-absorbed radiation onto a two-dimensional imaging device, or imager, which comprises an array of pixel detectors (simply called “pixels”). Such a system is shown in FIG. 1. Typically, the point source and the center of the two-dimensional imager lie on a common axis, which may be called the projection axis. The source's radiation emanates toward the imaging device in a volume of space defined by a right-circular, elliptical, or rectangular cone having its vertex at the point source and its base at the imaging device. For this reason, the radiation is often called cone-beam (CB) radiation. Generally, when no object is present within the cone, the distribution of radiation is substantially uniform on any circular area on the imager that is centered about the projection axis, and that is within the cone. However, the distribution of the radiation may be slightly non-uniform, but having rotational symmetry about the projection axis. In any event, any non-uniformity in the distribution can be measured in a calibration step and accounted for. The projection axis may not be at the center of the imager or the center of the object. It may pass through them at arbitrary locations including very near the edge.
In an ideal imaging system, rays of radiation travel along respective straight-line transmission paths from the source, through the object, and then to respective pixel detectors without generating scattered rays. However, in real systems, when a quantum of radiation is absorbed by a portion of the object, one or more scattered rays are often generated that deviate from the transmission path of the incident radiation. These scattered rays are often received by “surrounding” pixel detectors that are not located on the transmission path that the initial quantum of radiation was transmitted on, thereby creating errors in the electrical signals of the surrounding pixel detectors. In addition, background radiation from the environment is often present, which can be a function of time and the position of the two-dimensional imager, which is often rotated about the object during a scan. The source of the radiation is typically not a point but has some size. The effect of this is often characterized by a source distribution function. Rays from outside some area, usually a circle, of the source distribution function create spatial error which is not accounted for using typical data analysis techniques. Furthermore, in typical two-dimensional imagers, the radiation meant to be received by a pixel is often distributed by various components of the imager (e.g., scintillation plate), and received by surrounding pixels. Also, there is typically some electrical cross-talk in the electrical signals of the pixel detectors caused by the electrical circuitry that reads the array of pixel detectors. These two effects are often characterized by a point-spread function (PSF), which is a two-dimensional mapping of the amount of error caused in surrounding pixels by a given amount of radiation received by a central pixel. The surface of the PSF is similar to the flared shape of a trumpet output, with the greatest amount of error occurring in pixels adjacent to the central pixel. Each of these non-ideal effects creates spatial errors in the pixel data generated by the two-dimensional imager.
Also in an ideal radiation imaging system, each pixel detector outputs an electrical signal that is representative of the radiation that strikes its designated area on the imager, and that responds instantaneously to changes in received radiation without any lagging or memory effects. However, pixel arrays, which are typically constructed with a scintillation plate disposed over an array of semiconductor diodes, have a number of lagging effects (also called memory effects). The scintillation plate, which converts the radiation into light that can be readily detected by the semiconductor diodes, has a finite conversion-delay time where the plate's scintillation material continues to luminesce for a period of time after being struck by a quantum of radiation. Each semiconductor diode, which typically comprises amorphous silicon, generates pairs of free electrons and free holes in response to light received from the portion of the scintillation plate above it, which then produce an electrical signal as they travel to respective electrodes of the diode. However, semiconductor materials, particularly amorphous ones, have carrier traps that capture and hold free electrons and free holes in the semiconductor material for varying periods of time before releasing them. When so trapped, a carrier generated by the scintillation light does not generate a corresponding electrical current at the time it was generated, but rather generates an electrical current sometime afterwards, thereby causing a lag effect. The trapping of carriers also causes changes in signal detection gain and signal offset of the diode. Each of these non-ideal effects creates temporal errors in the pixel data generated by the two-dimensional imager.
As part of making their inventions, the inventors have found that the spatial errors and the temporal errors cause artifacts (e.g., phantom images) and loss of resolution and contrast in the CT image slices produced by the radiation imaging system. They have also found that these errors cause numerical errors in the image reconstruction algorithms (generally referred to as “CT number problems” in the art). All of the foregoing lead to image degradation. While some of the specific effects that give rise to the above-described spatial and temporal errors have been given, it may be appreciated that other effects may exist. Nonetheless, the present inventions also address the errors that arise from such other effects. As a general summary, spatial errors at a given pixel detector are caused by radiation that does not travel along the direct path from the radiation source to the pixel detector; these errors include radiation scattered by the object being imaged, radiation scattered by the components of the two-dimensional imager, and cross-talk in the electrical circuitry that reads the pixel array. Temporal errors are caused by radiation that was incident on the detector element prior to the current measurement, with the detection of the prior radiation being erroneously read in the current measurement because of delays in the scintillation material, semiconductor material, or other components. Accordingly, there is a need in the computerized tomography field to reduce the impacts of these spatial and temporal errors.